Control system for prosthetic knee

ABSTRACT

The invention relates to an automated speed-adaptive and patient-adaptive control scheme and system for a knee prosthesis. The control scheme and system utilizes sensory information measured local to the prosthesis to automatically adjust stance and swing phase knee resistances to a particular wearer under a wide variety of locomotory activities. Advantageously, no patient-specific information needs to be pre-programmed into the prosthetic knee by a prosthetist or the patient. The system is able to adapt to various types of disturbances once the patient leaves the prosthetist&#39;s facility because it is patient-adaptive and speed-adaptive.

RELATED APPLICATIONS

This application is a divisional of U.S. application Ser. No.10/646,097, filed Aug. 22, 2003, now U.S. Pat. No. 7,279,009 B2, issuedate Oct. 9, 2007, which is a continuation of U.S. application Ser. No.09/823,931, filed Mar. 29, 2001, now U.S. Pat. No. 6,610,101 B2, issuedate Aug. 26, 2003, which claims the benefit of U.S. ProvisionalApplication No. 60/192,966, filed Mar. 29, 2000, the entirety of eachone of which is hereby incorporated by reference herein.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to prosthetic knees in general and, inparticular, to a speed-adaptive and patient-adaptive control scheme andsystem for an external knee prosthesis.

2. Description of the Related Art

Most conventional active knee prostheses are variable torque brakeswhere joint damping is controlled by a microprocessor as an amputeewalks from step to step. Many brake technologies have been employed forknees including pneumatic, hydraulic and magnetorheological.

With most current prosthetic technology, a prosthetist adjusts kneeresistances to tune the artificial leg to the amputee so that the kneeprosthesis moves naturally at slow, moderate and fast walking speeds.During use, sensors local to the prosthesis are used to detect walkingspeed. A microprocessor then adjusts knee resistances based oncustomized values or data previously programmed by the prosthetist forthat specific patient only.

Disadvantageously, such a methodology for programming a prosthetic kneeis time consuming for both the prosthetist and the patient and has to berepeated for each patient. Moreover, any unforeseen changes in thepatient or in the patient's environment are not compensated for by theknee prosthesis after the patient has left the prosthetist's facility.This lack of adaptiveness in the knee system can disrupt normallocomotion and render the pre-programmed knee uncomfortable or evenunsafe. In this situation, the patient must return to the prosthetist'sfacility for the knee prosthesis to be reprogrammed. Again, undesirablythis results in additional wastage of time and further adds to the cost.

SUMMARY OF THE INVENTION

Accordingly it is one advantage of the present invention to overcomesome or all of the above limitations by providing an automatedspeed-adaptive and patient-adaptive control scheme and system for a kneeprosthesis. The control scheme and system utilizes sensory informationmeasured local to the prosthesis to automatically adjust stance andswing phase knee resistances to a particular wearer under a wide varietyof locomotory activities. Advantageously, no patient-specificinformation needs to be pre-programmed into the prosthetic knee by aprosthetist or the patient. The system is able to adapt to various typesof disturbances once the patient leaves the prosthetist's facilitybecause it is patient-adaptive and speed-adaptive.

In accordance with one preferred embodiment, a method is provided ofadaptively controlling the stance phase damping of a prosthetic kneeworn by a patient. The method comprises the step of providing a memoryin the prosthetic knee. The memory has stored therein correlationsbetween sensory data and stance phase damping established in clinicalinvestigations of amputees of varying body size. Instantaneous sensoryinformation is measured using sensors local to the prosthetic knee asthe patient stands, walks or runs. The instantaneous sensory informationis used in conjunction with the correlations to automatically adjuststance phase damping suitable for the patient without requiring patientspecific information to be pre-programmed in the prosthetic knee.

In accordance with another preferred embodiment, a method is provided ofadaptively controlling the swing phase damping torque of a prostheticknee worn by a patient as the patient travels at various locomotoryspeeds. The ground contact time of a prosthetic foot connected to theprosthetic knee by a prosthetic leg is indicative of the locomotoryspeed of the patient. The method comprises the step of continuouslymeasuring the contact time over periods of one gait cycle as the patientambulates at various locomotory speeds. The contact time is storedwithin a memory of the prosthetic knee in time slots corresponding tothe locomotory speed of the patient. The swing phase damping for kneeflexion is iteratively modulated to achieve a target peak flexion anglerange until the flexion damping converges within each time slot. Theswing phase damping for knee extension is iteratively modulated tocontrol the impact force of the extending prosthetic leg against anartificial knee cap of the prosthetic knee until the extension dampingconverges within each time slot. The converged damping values are usedto automatically control swing phase damping at all locomotory speeds.

In accordance with one preferred embodiment, an adaptive prosthetic kneeis provided for controlling the knee damping torque during stance phaseof an amputee. The prosthetic knee generally comprises a controllableknee actuator, sensors and a controller. The knee actuator provides avariable damping torque in response to command signals. The sensorsmeasure the force and moment applied to the prosthetic knee as theamputee moves over a supporting surface. The controller has a memory andis adapted to communicate command signals to the knee actuator andreceive input signals from the sensors. The memory has stored thereinrelationships between sensory data and stance phase damping establishedin prior clinical investigations of patients of varying body size. Thecontroller utilizes sensory data from the sensors in conjunction withthe relationships to adaptively and automatically control the dampingtorque provided by the knee actuator during stance phase independent ofany prior knowledge of the size of the amputee.

For purposes of summarizing the invention, certain aspects, advantagesand novel features of the invention have been described herein above. Ofcourse, it is to be understood that not necessarily all such advantagesmay be achieved in accordance with any particular embodiment of theinvention. Thus, the invention may be embodied or carried out in amanner that achieves or optimizes one advantage or group of advantagesas taught herein without necessarily achieving other advantages as maybe taught or suggested herein.

All of these embodiments are intended to be within the scope of theinvention herein disclosed. These and other embodiments of the presentinvention will become readily apparent to those skilled in the art fromthe following detailed description of the preferred embodiments havingreference to the attached figures, the invention not being limited toany particular preferred embodiment(s) disclosed.

BRIEF DESCRIPTION OF THE DRAWINGS

Having thus summarized the general nature of the invention and itsessential features and advantages, certain preferred embodiments andmodifications thereof will become apparent to those skilled in the artfrom the detailed description herein having reference to the figuresthat follow, of which:

FIG. 1 is a schematic drawing of one normal human locomotion cycleillustrating the various limb positions during stance and swing phases;

FIG. 2 is a schematic graphical representation of the variation in kneeangle showing state transitions during one normal gait cycle;

FIG. 3 is a plot of biological knee angle and mechanical power againstpercentage of a complete walking cycle for one subject;

FIG. 4 is a schematic illustration of a lower limb prosthetic assemblycomprising an electronically controlled prosthetic knee and havingfeatures and advantages in accordance with one preferred embodiment ofthe present invention;

FIG. 5 is a simplified block diagram representation of an adaptiveprosthetic knee system having features and advantages in accordance withone preferred embodiment of the present invention;

FIG. 6 is a diagram of one preferred embodiment of a state machinecontroller for the prosthetic knee system of FIG. 5 and showingstate-to-state transitional conditions for a gait or activity cycle;

FIG. 7 is a graph of foot contact time plotted against forward speed fora non-amputee moving at several steady state speeds;

FIG. 8 is a simplified schematic drawing illustrating the generaloverall configuration of one preferred embodiment of the prosthetic kneeactuator of the present invention;

FIG. 9 is a detailed exploded perspective view of a magnetorheologicallyactuated prosthetic knee brake having features and advantages inaccordance with one preferred embodiment of the present invention; and

FIG. 10 is a cross section view of the prosthetic knee of FIG. 9.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

In order for a trans-femoral (above-knee) amputee to walk in a varietyof circumstances, a prosthetic knee should provide stance control tolimit buckling when weight is applied to the limb. In addition, aprosthesis should provide swing phase control so that the knee reachesfull extension just prior to heel strike in a smooth and natural manner.

Unlike a biological knee, a prosthetic knee should accomplish bothstance and swing control without direct knowledge of its user's intentor of the environment. Rather, a prosthetic knee has to infer whetherthe amputee is walking, running, or sitting down. It should alsodetermine when subtle or drastic changes occur in the environment, suchas when the user lifts a suitcase or walks down a slope. Still further,the prosthesis should move naturally and be safe at all locomotoryspeeds, and should perform equally well for all amputees, independent ofbody weight, height, or activity level, without requiringpatient-specific information or programming from a prosthetist.

In accordance with one preferred embodiment of the present invention, aprosthetic knee is precisely and accurately controlled at substantiallyall locomotory speeds and for substantially all patients. The inventionutilizes an adaptation scheme that automatically adjusts stance andswing resistances or damping without pre-programmed information from apatient or prosthetist, making the “smart” knee both speed-adaptive andpatient-adaptive.

Normal Level-Ground Ambulation

Understanding normal human walking/running provides the basis for thedesign and development of effective lower limb prostheses withcontrolled motion. Normal human locomotion or gait can be described as aseries of rhythmical alternating movements of the limbs and trunk whichresult in the forward progression of the body's center of gravity.

One typical normal level-ground gait cycle, as schematically depicted inFIG. 1, comprises of the activity that occurs between heel strike of onelower limb I0 and the subsequent heel strike of the same limb 10. Thelimb or leg 10 generally comprises a foot 12 and a shin or shank portion14 coupled or articulated to a thigh portion 16 via a knee or knee joint18. During a single gait cycle each lower limb or extremity passesthrough one stance or extended phase 20 and one swing phase 22.

The stance phase 20 begins at heel-strike 24 when the heel touches thefloor or supporting ground surface and the stance knee begins to flexslightly. This flexion allows for shock absorption upon impact and alsomaintains the body's center of gravity at a more constant vertical levelduring stance.

Shortly after heel-strike 24, the sole makes contact with the ground atthe beginning of the foot-flat phase 26. After maximum flexion isreached in the stance knee, the joint begins to extend again, untilmaximum extension is reached at mid-stance 28 as the body weight isswung directly over the supporting extremity and continues to rotateover the foot.

As the body mass above the ankle continues to rotate forward, the heellifts off the ground at heel-off 30. Shortly after this, the body ispropelled forward by the forceful action of the calf-muscles (poweredplantar-flexion). The powered plantar-flexion phase terminates when theentire foot rises from the ground at toe-off 32.

During late stance, the knee of the supporting leg flexes in preparationfor the foot leaving the ground for swing. This is typically referred toin the literature as “knee break”. At this time, the adjacent footstrikes the ground and the body is in “double support mode”, that is,both the legs are supporting the body weight.

At toe-off 32, as the hip is flexed and the knee reaches a certain angleat knee break, the foot leaves the ground and the knee continues to flexinto the swing phase. During early swing the foot accelerates. Afterreaching maximum flexion at mid-swing 34, the knee begins to extend andthe foot decelerates. After the knee has reached full extension, thefoot once again is placed on the ground at heel-strike 24′ and the nextwalking cycle begins.

Typically, the anatomical position is the upright position, thereforeflexion is a movement of a body part away from the extended or stance oranatomical position. Thus, bending of the knee is knee flexion.Extension is a movement of a limb towards the anatomical position, thusknee extension is a movement in the “straightening” direction.

Stated differently, if a knee joint is looked at as a simple hinge,there are two separate actions which can occur. In “flexion”, the kneejoint rotates to enable the upper and lower leg segments to move closertogether. In “extension” the knee joint rotates in the oppositedirection, the leg segments move apart and the leg straightens.

During a typical normal walking progression on a generally levelsurface, the maximum flexion angle α_(F) varies between about 60° and80°. The maximum extension angle α_(E) is typically about or close to180°. Thus, in level walking the normal human knee rotates through arange of approximately 60°-80° going from a position of full extensionin late stance to 60°-80° of flexion shortly after toe-off. In othersituations, for example, in a sitting position, the maximum flexionangle α_(F) can be about 140°-150°.

Referring to FIG. 2, preferably, the gait cycle of FIG. 1 is categorizedinto five distinct states or phases. FIG. 2 schematically shows the kneeangle θ, that is, the angle the knee rotates from full extension, withstate or phase transitions during activity that occurs between the heelstrike (HS) of one lower limb and the subsequent heel strike (HS) of thesame limb. The x-axis 36 represents time between consecutive heelstrikes of the walking cycle. The y-axis 38 represents the knee angle θ.

State 1 represents early stance flexion just after heel strike (HS).State 2 represents early or mid stance extension after maximum stanceflexion is reached in State 1. State 3, or knee break, typically occursat the end of stance, beginning just after the knee has fully extendedand terminates when the foot has left the ground at toe-off (TO). State4 represents a period of knee flexion during the swing phase of awalking or running cycle. State 5 represents a period of knee extensionduring the swing phase of a walking or running cycle, after maximumswing flexion is reached in State 4.

As discussed later herein, these basic states or phases suggest theframework of a prosthetic knee controller as a state machine. Thus, FIG.2 is a graphical representation of a person moving through a normal gaitcycle and the location of each state within that cycle. Table 1 belowsummarizes the activity during each of the States 1 to 5.

TABLE 1 State Activity 1 Stance Flexion 2 Stance Extension 3 Knee Break4 Swing Flexion 5 Swing Extension

FIG. 3 is a plot of typical biological knee angle and knee power versustime normalized to the step period (adapted from Grimes, 1979). Thex-axis 40 represents time normalized to the step period, T, orpercentage of walking cycle. The y-axis 42 represents knee power (P inft-lb/sec) and the y-axis 44 represents knee angle (θ in degrees).

In FIG. 3, four walking trials are shown for one subject. Zero percentand one hundred percent mark two consecutive heel strikes of the sameleg and zero angle generally corresponds to the heel strike angle. Also,in FIG. 3, RHS represents right heel strike, RFF represents right flatfoot, LTO represents left toeoff, RHO represents right heel off. LHSrepresents left heel strike, LFF represents left flat foot, RTOrepresents right toe off and LHO represents left heel off.

Still referring to FIG. 3, the smaller dip 46 in the angle plot (about15% of the full cycle) represents the flexion and extension of the kneeduring early or mid stance, whereas the larger dip 48 (about 75% of thefull cycle) represents the flexion and extension of the knee during theswing phase. Throughout the cycle, the knee mechanical power isprimarily negative or dissipative. This justifies the use or employmentof a variable damper or a variable torque brake in a knee prosthesis.Such a variable damper or knee actuator is discussed further hereinbelow.

System Configuration

FIG. 4 is a schematic illustration of a lower limb prosthetic assemblyor prosthesis 100 comprising an electronically controlled active kneeprosthesis 110 and having features and advantages in accordance with onepreferred embodiment of the present invention. As described in greaterdetail later herein, preferably, the active knee prosthesis comprises avariable-torque braking system or damper 130 and an onboard control unitor system 120 housed in a supporting frame 141. The prosthetic kneesystem 110 provides resistive forces to substantially simulate theposition and motion of a natural knee joint during ambulation and/orother locomotory activities performed by the amputee.

At one end the artificial knee system 110 is coupled or mechanicallyconnected to a residual limb socket 102 which receives a residual limb,stump or femur portion 104 of the amputee. The other end of theprosthetic knee 110 is coupled or mechanically connected to a pylon,shin or shank portion 106 which in turn is coupled or mechanicallyconnected to a prosthetic or artificial foot 108.

Advantageously, the prosthetic knee system 110 of the preferredembodiments is both speed-adaptive and patient-adaptive. Thus, the kneejoint rotation is automatically controlled at substantially all speedsand for substantially all patients, regardless of body size, withoutpre-programmed information or calibrated data from a patient orprosthetist.

One main advantage of the preferred embodiments of the knee system isthat it is able to adapt to various types of disturbances once thepatient leaves the prosthetist's facility because it is patient-adaptiveand speed-adaptive. As an example, when the patient picks up a suitcase,the knee responds to the disturbance automatically. With conventionaltechnology, the patient would have to go back to the prosthetistfacility to re-program their knee. In the preferred embodiments, thetrial period is not typically “lengthy” and “fatiguing”.

The prosthetic knee 110 of the preferred embodiments advantageouslypermits the amputee to move and/or adapt comfortably and safely in awide variety of circumstances. For example, during walking, running,sitting down, or when encountering subtle or drastic changes in theenvironment or ambient conditions, such as, when the user lifts asuitcase or walks down a slope.

The artificial knee 110 provides stance control to limit buckling whenweight is applied to the limb. In addition, the prosthetic knee 110provides aerial swing control so that the knee reaches full extensionjust prior to or at heel-strike in a smooth and natural manner.

Preferably, the artificial knee system 110 of the present invention isused in conjunction with a trans-femoral (above-knee, A/N) amputee.Alternatively or optionally, the prosthetic knee 110 may be adapted foruse with a knee-disarticulation (K/D) amputee where the amputation isthrough the knee joint, as needed or desired, giving due considerationto the goals of achieving one or more of the benefits and advantages astaught or suggested herein.

Knee Electronics

FIG. 5 illustrates one preferred embodiment of the prosthetic kneesystem 110 of the invention in block diagram format. In FIG. 5, thesolid communication lines represent signal/data flow and the phantom ordashed communication lines represent energy flow.

As stated above, preferably, the automated prosthetic knee system 110generally comprises a variable-torque braking system or damper 130 andan onboard control unit or system 120. The feedback control system 120comprises a central controller 132 which receives sensory and diagnosticinformation to control the operation of the knee actuator 130 and otherassociated equipment (as discussed below). For purposes of clarity, thevarious components of the prosthetic knee system 110, in accordance withone preferred embodiment, are listed in Table 2 below.

TABLE 2 Component(s) Reference Numeral Knee Actuator 130 Microprocessor132 Knee Angle Sensor 134 Knee Angle Amplifier 136 Knee AngleDifferentiator 138 Axial Force and Moment Sensors 140 Axial Force andMoment Amplifiers 142 Battery Monitoring Circuit 144 Moisture DetectionCircuit 146 Power Usage Monitoring Circuit 148 Memory 150 SerialCommunications Port 152 Safety Mechanism 154 Safety Mechanism Driver 156Safety Watchdog Circuit 158 Knee Actuator Current Amplifier 160 AudibleWarning Transducer 162 Audible Warning Circuit 164 Vibration Transducer166 Vibration Warning Generator 168 Battery 170 Battery ProtectionCircuitry 172 Battery Charge Circuit 174 Circuit Power Supplies 176Circuit Power Conditioners 178

As mentioned above, the knee actuator 130 comprises a variable torquebrake or damper for modulating joint damping to control extension andflexion movements based on command signals from the knee controller 132.The manner in which the control scheme of the preferred embodimentscontrols knee joint rotation is discussed in further detail laterherein.

The knee actuator or brake 130 can comprise any one of a number ofconventional brakes. These include without limitation (i) dry frictionbrakes where one material surface rubs against another surface withvariable force; (ii) viscous torque brakes using hydraulic fluidsqueezed through a variable sized orifice or flow restriction plate; and(iii) magnetorheological (MR) brakes or dampers where MR fluid(containing small iron particles suspended in the fluid) is squeezedthrough a fixed orifice or flow restriction plate, with viscosity of thefluid being varied in response to an applied magnetic field. Optionally,the knee brake 130 comprises a pneumatic brake, as known in the art.

In one preferred embodiment, and as discussed in further detail laterherein, the knee brake 130 comprises a variable torque rotarymagnetorheological (MR) brake that operates in the shear mode. MR fluidis sheared between a plurality of rotors and stators to generate avariable and controlled damping effect which precisely and accuratelymodulates the knee joint rotation.

In one preferred embodiment, the prosthetic knee system 110 comprises anartificial knee cap or extension stop to limit the maximum kneeextension. The artificial or prosthetic knee cap is preferably below theknee actuator 130 and is mechanically connected to the knee actuator 130and/or the frame 141.

The knee actuator current amplifier 160 comprises a circuit whichgenerates the needed or desired current from the battery 170 in the kneeactuator 130 to modulate the damping torque provided by the knee brake130. Command signals or instructions from the microprocessor 132 to theknee actuator current amplifier 160 determine the current supplied tothe knee actuator 130, and hence the amount of damping torque generated.

The onboard microprocessor 132 including memory 150 are local to theprosthetic knee system 110. The microprocessor 132 is the primarycomputational unit of the prosthetic knee system 110 and receives inputelectrical signals from the various components of the knee system 110,processes them, and generates output electrical signals to monitor andcontrol the actuations of the prosthetic knee 130 and other associatedcomponents, as necessary.

The microprocessor 132 includes circuitry which digitizes incomingsignals and generates outgoing analog signals. The microprocessorfurther includes timing modules and watchdog self-resetting circuitry.The memory 150 comprises internal or external volatile and non-volatilememory.

The microprocessor 132 preferably comprises a Motorola 68HC12B32 16 bitseries microprocessor. This processor has 8 channel analog to digitalconversion capability, 32K of flash and 768 bytes of EEProm memory. Theexternal memory comprises two industry standard 32K by 8 bit staticRAMs. The serial flash is an Atmel AT45D081 and uses the serialcommunications interface (SCI) provided by the microprocessor.

The serial communications port 152 provides an interface between theknee electronics, via the microprocessor 132, and external diagnostic,data logging and programming equipment. The port 152 can efficaciouslycomprise any one of a number of commercially available communicationports, for example, RS232, RS485, ethernet and the like, as needed ordesired, giving due consideration to the goals of achieving one or moreof the benefits and advantages as taught or suggested herein.

The microprocessor 132 along with the other associated sensory,diagnostic safety and protection circuitry of the prosthetic knee system110 are preferably mounted on a circuit board to provide a compactassembly. The circuit board is preferably housed within and secured tothe frame 141 directly or utilizing an intermediate shell or cover toprotect the circuit board and components mounted thereon.

The knee angle sensor 134 is used to encode the absolute knee angle.Preferably, the knee angle sensor 134 measures the degree to which asingle degree-of-freedom knee joint is flexed or extended. The kneeangle amplifier 136 comprises a circuit which conditions the signalreceived from the knee angle sensor 134 and feeds it to themicroprocessor 132 for knee control purposes, as discussed below.

The knee angle differentiator 138 comprises a circuit whichdifferentiates the signal received from the knee angle sensor 134 todetermine the rotational or angular velocity of the knee and feeds thissignal to the microprocessor 132 for knee control purposes, as discussedbelow. The knee angular velocity signal further determines whether theknee is flexing or extending.

The angle sensor 134 is preferably mounted on the frame 141 (FIG. 4).Alternatively, the angle sensor 134 is mounted on the side of the kneeactuator 130 or directly below the knee actuator 130, as needed ordesired.

In one preferred embodiment, the angle sensor 134 comprises an anglesensing potentiometer. In another preferred embodiment, the angle sensor134 comprises an optical shaft encoder. In yet another preferredembodiment, the angle sensor 134 comprises a magnetic shaft encoder. Inother preferred embodiments, alternate knee angle sensing devices may beutilized with efficacy, as required or desired, giving due considerationto the goals of accurately estimating the knee angle, and/or ofachieving one or more of the benefits and advantages as taught orsuggested herein.

The axial force and moment sensors 140 comprise a transducer thatgenerates signals proportional to the lower leg axial force and momentor torque. In one preferred embodiment, the transducer comprises a forestrain gage sensor and an aft strain gage sensor. To compute axialforce, the fore and aft signals are added, and to compute the moment,the signals are subtracted. The axial force and moment amplifiers 142condition the signals received from the axial force and moment sensors140 and feed it to the microprocessor 132 for knee control purposes, asdiscussed below.

The axial force sensors 140 measure the component of force applied tothe knee prosthesis 110 from the ground or other supporting surface in adirection substantially along or parallel to the shin longitudinal axis180 (FIG. 4) or knee long axis. The axial force measurement is used todetermine whether the prosthetic foot 108 (FIG. 4) is on or off theground or other supporting surface. That is, a zero axial forceindicates that the foot 108 is not on the ground, for example, in theswing phase, while a non-zero axial force indicates that the foot 108 ison the ground, for example, in the stance phase.

The torque or moment sensors 140 measure the component of torque appliedto the knee prosthesis 110 in a medial-lateral direction 182 as shown inFIG. 4. In addition, the moment sensors 140 determine whether theapplied knee moment is a flexion or extension moment. Typically, at heelstrike a flexion moment is applied to the knee prosthesis 110, tendingto flex the knee joint, and throughout late stance an extension momentis applied, tending to extend the joint.

The axial force and moment sensors 140 are preferably mounted on theframe 141 (FIG. 4). In one preferred embodiment, the axial force andmoment sensors 140 comprises a strain gauge load cell. In anotherpreferred embodiment, the axial force and moment sensors 140 comprise adeflection encoded shock/spring mechanism. In other preferredembodiments alternate load and/or moment sensing devices may be utilizedwith efficacy, as required or desired, giving due consideration to thegoals of accurately estimating the axial load and/or applied moment,and/or of achieving one or more of the benefits and advantages as taughtor suggested herein.

In one preferred embodiment, the axial force and moment sensors 140comprise a plurality of strain gauges. Preferably, four gauges are usedwith two strain gauges mounted on the front 184 of the frame 141 and twostrain gauges mounted on the rear 186 of the frame 141 to measure anddifferentiate between load on the heel of the foot 108 and load on thetoe of the foot 108. Stated otherwise, the strain measurement providesan indication as to whether the center of gravity is in an anterior,centered or posterior position relative to the prosthetic foot 108.

The strain gauges are preferably arranged in a wheatstone bridgeconfiguration to generate an electric signal which changesproportionally with bending moment strain. As the skilled artisan willrecognize, such a wheatstone bridge configuration is a standardarrangement for determining the resistance change of strain gauges.

The battery monitoring circuit 144 continuously or periodically monitorsthe battery voltage, current and temperature for safety purposes. Thedata from the battery monitoring circuit 144 is continuously orperiodically provided to the microprocessor 132 to facilitate inconstraining the knee operation to within the battery manufacturer'sspecification.

The moisture detection circuit 146 continuously or periodically monitorsthe moisture levels for safety purposes and senses any abnormal moistureon the system circuit board and/or other associated system circuitry dueto condensation, submersion and the like. The data from the moisturedetection circuit 146 is continuously or periodically provided to themicroprocessor 132.

In one preferred embodiment, the moisture detection circuit 146comprises interdigitated copper traces. In other preferred embodiments,the moisture detection circuit can comprise alternate moisture detectingdevices with efficacy, as required or desired, giving due considerationto the goals of reliably detecting moisture levels on the systemelectronics, and/or of achieving one or more of the benefits andadvantages as taught or suggested herein.

The power usage monitoring circuit 148 continuously or periodicallymeasures the power consumption by the knee actuator 130 for safetypurposes. The data from the power usage monitoring circuit 148 iscontinuously or periodically provided to the microprocessor. Inaddition, the power usage monitoring circuit 148 or other independentcircuits may be utilized, as needed or desired, to measure the powerconsumption by other electronic components of the prosthetic knee system110.

The prosthetic knee system 110 preferably comprises a safety systemincluding the safety mechanism 154. The safety mechanism 154 is actuatedor activated to put the system 110 in a default safety mode when asystem error is detected by the microprocessor 132. Such a system errorcan occur if abnormal behavior is noted in any of the signals from theknee angle sensors 134, the axial force and moment sensors 140, thebattery monitoring circuit 144, the moisture detection circuit 146 andthe power usage monitoring circuit 148 indicating a system malfunctionand/or other concern over the integrity of the knee actuator 130.

Detection of a system error causes the safety mechanism or actuator 154to activate a safety default mode such that even with a systemmalfunction the prosthetic knee system 110 remains safe for the amputee.For example, in the safety default mode, the knee could resist flexionbut could be free to extend, thereby ensuring the safety of the patient.

The safety mechanism driver 156 comprises a power amplifier that turnson or off the safety default mode of the safety mechanism 154 based oncommand signals or instructions from the microprocessor 132. The safetywatchdog circuit 158 comprises a circuit which is periodically orcontinuously “attended” to by signals from the microprocessor 132 toprevent the watchdog circuit 158 from unnecessarily enabling the safetydefault mode by sending signals to the safety mechanism driver 156. Inother words, the safety watchdog circuit 158 would activate the safetymechanism 154 unless otherwise periodically or continuously instructedso by the microprocessor.

Preferably, and when possible, to warn the user of a system malfunctionor unusual operating condition, prior to the activation of the defaultsafety mode, either one or both of the audible warning transducer 162and the vibration transducer 166 are activated. The audible warningcircuit 164 comprises an amplifier which generates an electronic signalto create audible noise by the warning transducer 162 when enabled. Theaudible warning circuit 164 receives command signals or instructionsfrom the microprocessor 132.

The audible warning transducer 162 is preferably housed in or secured tothe frame 141 (FIG. 4). In one preferred embodiment, the audible warningtransducer 162 comprises a piezo speaker. In other preferredembodiments, alternate sound generating devices may be utilized withefficacy, as required or desired, giving due consideration to the goalsof warning the user, and/or of achieving one or more of the benefits andadvantages as taught or suggested herein.

The vibration transducer 166 comprises an actuator which vibrates theprosthetic knee system 110 in such a way as to draw attention from thewearer. The vibration warning generator 168 comprises an amplifier whichgenerates an electronic signal to turn on the vibration transducer 164when enabled. The vibration warning generator 168 receives commandsignals or instructions from the microprocessor 132.

The vibration transducer 166 is preferably mounted on the system circuitboard. Alternatively, the vibration transducer 166 is housed in orsecured to the frame 141 (FIG. 4). In one preferred embodiment, thevibration transducer 166 comprises a wobble motor. In other preferredembodiments, alternate vibration generating devices may be utilized withefficacy, as required or desired, giving due consideration to the goalsof warning the user, and/or of achieving one or more of the benefits andadvantages as taught or suggested herein.

The onboard battery or power source 170 supplies power to the kneeactuator 130, the safety mechanism 154, the audible warning transducer162 and the vibration transducer 166. The circuit power conditioners 178convert the raw battery power to power that is conditioned for use bythe microprocessor 132 and other sensory circuitry and individual systemsubcircuits. The circuit power supplies 176 provide the conditionedpower to the microprocessor 132 and other sensory circuitry andindividual system subcircuits.

Thus, via the circuit power supplies 176 and the circuit powerconditioners 178, the battery 170 distributes power to themicroprocessor 132 and other sensory circuitry and individual systemsubcircuits including the knee angle amplifier 136, the knee angledifferentiator 138, the axial force and moment amplifiers 142, thebattery monitoring circuit 144, the moisture detection circuit 146, thepower usage monitoring circuit 148, the safety watchdog circuit 158, thesafety mechanism driver 156, the knee actuator current amplifier 160,the audible warning circuit 164, the vibrator warning generator 168 andany other associated circuits, as necessary.

The battery protection circuitry 172 protects the battery 170 fromexceeding safe operating conditions. If desired, a battery state ofcharge indicator may also be provided. The battery charge circuitry 174converts power from a charging source, typically a wall outlet, to thepower levels suited for the battery 170.

The Control Scheme

The State Machine

The basic phases or states of biological gait (as discussed above)suggest the framework of the prosthetic knee controller as a statemachine. Thus, each phase corresponds to a State 1 to 5 (see, forexample, FIG. 2 and Table 1). FIG. 6 is a diagram of one preferredembodiment of a state machine controller 190 of the prosthetic kneesystem 110 and shows state-to-state transitional conditions.

As discussed above, the onboard knee angle sensor 134 measures the kneeangle and the onboard axial force and moment sensors 140 measure theaxial force and the knee moment. The knee angle data, the kneerotational velocity data, the axial force data and the knee moment dataare provided to the microprocessor or main controller 132 to determinesystem state, and accordingly automatically control the actuations ofthe knee brake or actuator 130 to modulate knee joint rotation.

Also as discussed above, the knee angle signal determines the degree ofknee joint rotation and the knee angular velocity signal determineswhether the knee is flexing or extending. The axial force measurementdetermines whether the prosthetic foot is on or off the ground or othersupporting surface. The knee moment measurement determines whether theapplied knee moment is a flexion or extension moment.

Based upon these sensory data provided to the microprocessor 132, thestate machine controller 190 cycles through the various States 1, 2, 3,4 and 5 as the user moves through each gait cycle or other locomotoryactivity. Often, and as seen in FIG. 6, the controller 190 changes statedepending on whether the moment is above or below an extension momentthreshold or critical value. Advantageously, and as discussed below,these threshold moments are automatically self-learned or self-taught bythe prosthetic knee system of the preferred embodiments for eachindividual patient without pre-programmed information about the specificpatient.

Preferably, the control of the state machine 190 on the behavior of theknee damper 130 allows the patient to perform a wide variety ofactivities. These include normal walking or running on a level orinclined surface, sitting down, ascending or descending steps or othersituations, for example, when a user lifts a suitcase. Again, in theseand other situations, the prosthetic knee system of the preferredembodiments automatically provides for accurate knee damping controlwithout pre-programmed information about the specific patient.

The overall operation of the state machine controller 190 and thevarious conditions that are satisfied between state-to-state transitionsare now described in accordance with one preferred embodiment. Based onthe input sensory data (as described above) these provide information tothe knee brake 130 on how to modulate knee damping. The control actionsfor each state are described later herein.

First, the state transitions and conditions for these transitions aredescribed for a typical walking or running cycle. As stated above, theaxial force is the component of force applied to the knee prosthesis 110from the ground or other supporting surface in a direction substantiallyalong or parallel to the shin longitudinal axis 180 (FIG. 4) or kneelong axis. The applied moment is the component of torque applied to theknee prosthesis 110 in a medial-lateral direction 182 as shown in FIG.4.

State 1 (stance flexion) transitions to State 2 (stance extension) undercondition C12. Condition C12 is satisfied when the knee first achieves asmall extension velocity. At this stage, the prosthetic foot is on theground or other supporting surface.

State 2 (stance extension) transitions to State 3 (knee break) underconditions C23. Conditions C23 are satisfied when the extension momentis below a threshold or critical level or value, when the knee is at orclose to full extension, and when the knee has been still for a certainamount of time.

State 3 (knee break) transitions to State 4 (swing flexion) undercondition C34. Condition C34 is satisfied when the axial force fallsbelow a threshold or critical level or value. That is, at this stage theprosthetic foot is off or nearly off the ground or other supportingsurface.

State 4 (swing flexion) transitions to State 5 (swing extension) undercondition C45. Condition C45 is satisfied when the knee first begins toextend. At this stage, the prosthetic foot is still off the ground orother supporting surface.

State 5 (swing extension) transitions back to State 1 (stance flexion)under condition C51. Condition C51 is satisfied when the axial forceclimbs above a threshold or critical level or value. This completes onewalking or running gait cycle.

As indicated above, the state-to-state transitions may follow otherpatterns than the State 1 to State 2 to State 3 to State 4 to State 5scheme depending on the particular activity of the amputee and/or theambient or terrain conditions. Advantageously, the finite state machinecontroller 190 automatically adapts to or accommodates for situations inwhich alternate state transitions may occur to provide the amputee withoptions of achieving a wide variety of substantially life-like ornatural movements under diverse external conditions.

State 1 (stance flexion) transitions to State 3 (knee break) underconditions C13. Conditions C13 are satisfied when the extension momentis below a threshold or critical level or value, when the knee is at orclose to full extension, and when the knee has been still for a certainamount of time. This state transition from State 1 to State 3 can occurwhen the amputee fails to go through the normal flexion-extension cycleduring stance.

State 1 (stance flexion) transitions to State 4 (swing flexion) undercondition C14. Condition C14 is satisfied when the axial force fallsbelow a small but nonzero threshold or critical level or value. Thisstate transition from State 1 to State 4 can occur when the amputeestands on the knee but alternately shifts back and forth, weighting andunweighting the prosthesis.

State 2 (stance extension) transitions to State 1 (stance flexion) undercondition C21. Condition C21 is satisfied when the knee achieves a smallbut nonzero flexion velocity. This state transition from State 2 toState 1 can occur if the amputee begins to flex the knee during theextension period of stance.

State 2 (stance extension) transitions to State 4 (swing flexion) undercondition C24. Condition C14 is satisfied when the axial force fallsbelow a threshold or critical level or value. This state transition fromState 2 to State 4 can occur if the amputee lifts his foot during theextension period of stance.

State 3 (knee break) transitions to State 1 (stance flexion) underconditions C31. Conditions C31 are satisfied when the knee has been inState 3 for a certain amount of time, OR if the extension moment isabove a threshold or critical level AND when the knee is at fullextension or close to full extension. This state transition from State 3to State 1 can occur if the amputee leans back on his heels from astanding position.

State 4 (swing flexion) transitions to State 1 (stance flexion) undercondition C41. Condition C41 is satisfied when the axial force climbsabove a small but nonzero threshold or critical value. This statetransition from State 4 to State 1 can occur if the amputee stands onthe knee but alternately shifts back and forth, weighting andunweighting his prosthesis.

As discussed above, based upon input sensory data, the controller 190cycles through the states as the user moves through each gait cycle oractivity. The state machine software is resident within themicroprocessor 132 or memory 150. Next, the various control actions orscheme for each state are described. The control scheme for States 1, 2and 3 is referred to as “stance phase control” and the control schemefor States 4 and 5 is referred to as “swing phase control.”

Stance Phase Control

In accordance with one preferred embodiment, a scheme is provided toadaptively control the stance phase damping of a prosthetic knee worn bya patient. Stored in the memory of the prosthetic knee are correlationsrelating sensory data and stance phase damping. Established in clinicalinvestigations of amputees of varying body size these relationscharacterize knee behavior when the prosthetic foot is in contact withthe ground. Sensory information are used in conjunction with thesecorrelations to define how stance phase damping should be modulated instanding, walking and running.

In accordance with one preferred embodiment, an adaptive prosthetic kneeis provided for controlling the knee damping torque during stance phaseof an amputee. The prosthetic knee generally comprises a controllableknee brake, sensors and a controller. The knee brake provides a variabledamping torque in response to command signals. The sensors measure kneeangle, axial force and applied moment as the amputee moves over asupporting surface. The controller has a memory and is adapted tocommunicate command signals to the knee brake and receive input signalsfrom the sensors. The memory has stored therein relationships betweensensory data and stance phase damping established in prior clinicalinvestigations of patients of varying body size. In addition,biomechanical information is stored in memory to guide the modulation ofdamping profiles. The controller utilizes sensory data from the sensorsin conjunction with both clinical and biomechanical information toadaptively and automatically control the damping torque provided by theknee brake during stance phase independent of any prior knowledge ofpatient size.

State 1 (Stance Flexion) and State 2 (Stance Extension):

In normal gait, the knee first flexes and then extends throughout earlyto midstance (see FIGS. 2 and 3). In State 1, or stance flexion, aprosthetic knee should preferably exert a resistive torque or damping toinhibit the knee from buckling under the user's weight. A prostheticknee should also preferably exert a resistive torque or damping duringthe extension period of stance, or State 2, to slow or damp kneeextension so that the knee does not overextend, thereby preventing therotating portion of a knee, such as the knee brake, to slam against aprosthetic kneecap (extension stop) or outer knee cover.

The degree to which a prosthetic knee should dampen flexion andextension so as to closely simulate a life-like or natural response islargely dependent on body weight. That is, in States 1 and 2 largerdamping values are preferred for larger users so as to more faithfullysimulate a generally life-like or natural feel. (Note that in general atall user does require a greater knee resistance but tall peopletypically tend to rotate the knee faster thereby increasing the torqueresponse of the system—current is proportional to knee rotationalvelocity where the proportionality constant is knee damping.)

In accordance with one preferred embodiment, clinical studies wereperformed with amputees of different body sizes ranging from small/lightto large/heavy to generally capture the full range of body sizes. Theseusers utilized prosthetic knees and other sensory equipment. Preferably,the users utilized the prosthetic knee brake 130 along with the axialand moment sensors 140 and the knee angle sensor 134.

In these clinical investigations, flexion and extension damping valuesprovided by the knee actuator 130 were optimized for amputees ofdifferent body size while monitoring the axial force, knee moment, kneeangle and knee angular velocity data, among other associated data, asnecessary. These data were then used to establish relationships orcorrelations between stance phase resistances and sensory informationmeasured and/or computed during stance.

Preferably, the clinical study data is collected over a wide variety ofpatient activities and/or external conditions and terrain. These includenormal walking or running on a level or inclined surface, sitting down,ascending or descending steps or other situations, for example, when auser lifts a suitcase, among other.

The optimized stance phase knee resistance or damping and sensory datarelationships or correlations for patients of varying body size arestored or programmed in the controller or microprocessor 132 or systemmemory 150. These are used in the prosthetic knee system 110 of thepreferred embodiments to automatically control the actuations of theknee brake 130.

When an amputee first walks utilizing the prosthetic knee system 110 ascontrolled by the preferred control schemes of the invention,preferably, the microprocessor or controller 132 initially sets State 1damping or resistance to knee rotation to a large value. For a lineardamper in which torque is proportional to knee rotational velocity, anadequate proportionality constant, or damping value, is 20 Nm*secondsper radian. This ensures that the prosthetic knee 110 is safe and doesnot buckle to exceedingly large flexion angles. Preferably, this maximumflexion angle does not exceed 15°.

In distinction to initial State 1 damping, preferably, themicroprocessor or controller 132 initially sets State 2 damping orresistance to knee rotation to a smaller value. For a linear damper inwhich torque is proportional to knee rotational velocity, an adequateproportionality constant, or damping value, is 10 Nm*seconds per radian.This allows the amputee to extend the knee even if the knee happens tobecome flexed.

As the amputee starts moving and taking several steps, the axial forceand moment sensors 140 and the angle sensor 134 are continuously orperiodically providing axial force, applied moment, knee angle and kneeangular velocity data or signals to the microprocessor or controller132. These sensory data, and in particular the peak force and peaktorque and/or the axial force and torque profiles applied to theprosthetic knee system 110, are used by the controller 132 to adjust theflexion and extension damping to values or profiles that were determinedto give reasonable or optimized or generally life-like stance behaviorduring the prior clinical investigations.

As discussed above, the relationships or correlations obtained duringthese clinical investigations of a wide range of patients having varyingbody sizes have been programmed or stored in the controller 132. As thepatient continues to use the prosthetic knee system 110, furtherautomated refinements and fine-tuning can be made by the system 110, asnecessary.

The prosthesis of the preferred embodiments is a self-teaching and/orself-learning system that is guided by clinical (prosthetic) andbiomechanical knowledge. For example, biomechanical knowledge (stored inthe system memory) includes information related to the mechanics oftypical human walking/running, as discussed above in reference to FIG.1.

Moreover, the clinical relationships or correlations also allow theprosthetic knee system 110 to determine the appropriate “thresholdmoments” for the particular amputee independent of body size. Asdiscussed above, these threshold moments are used by the state machine190 (FIG. 6) to change state depending on whether the threshold momentis above or below certain values specific to the patient.

Advantageously, in the preferred embodiments, no patient-specificinformation needs to be pre-programmed into the prosthetic knee by aprosthetist or the patient. Using sensory information measured local tothe knee prosthesis, stance resistances automatically adapt to the needsof the amputee, thereby providing an automated patient-adaptive system.

State 3 (Knee Break):

In one preferred embodiment, State 3 (knee break) knee damping orresistance is maintained substantially constant and minimized so thatthe amputee can easily flex the knee. Preferably, this minimum value ofthe knee damping torque is about 0.4 N-m and is largely determined bythe particular knee brake utilized. Alternatively, other minimum dampingtorque values and/or variable torques may be utilized with efficacy, asneeded or desired, giving due consideration to the goals of achievingone or more of the benefits and advantages as taught or suggestedherein.

In another preferred embodiment, the State 3 knee damping or torque isdetermined as described above for States 1 and 2. That is, measuredsensory data, and in particular the peak force and peak torque and/orthe axial force and torque profiles applied to the prosthetic kneesystem 110, are used by the controller 132 to adjust the knee resistanceor damping to values or profiles that were determined to give reasonableor optimized or generally life-like stance behavior during priorclinical investigations.

Swing Phase Control

In accordance with one preferred embodiment, a scheme is provided ofadaptively controlling the swing phase damping torque of a prostheticknee worn by a patient as the patient travels at various locomotoryspeeds. The ground contact time of a prosthetic foot, measured from heelstrike to toe-off, has been shown to correlate well with forwardlocomotory speed. The scheme comprises the step of continuouslymeasuring foot contact time as an estimate of the patient's forwardspeed, and adaptively modulating swing phase damping profiles until theknee is comfortable and moves naturally. The swing phase damping profilefor knee flexion is iteratively modulated to achieve a particular rangeof peak flexion angle. In distinction, for knee extension, knee dampingis modulated to control the impact force of the extending leg againstthe artificial knee cap. The converged damping values are used toautomatically control swing phase damping at all locomotory speeds.

In one preferred embodiment, during stance phase the controller 132computes a parameter, based on input sensory data, that changes withlocomotory speed of the amputee. Preferably, this parameter changesmonotonically with locomotory speed. As discussed below, this parameteris used by the controller 132 to automatically control swing phase kneeresistances for substantially all patients at substantially all speeds.

In one preferred embodiment, the speed control parameter is the amountof time the prosthetic foot remains in contact with the ground, or footcontact time. In another preferred embodiment, the speed controlparameter is the maximum flexion velocity that occurs betweensubstantially maximum or full extension and about thirty degrees flexionas the leg prosthesis flexes from State 3 to State 4. In other preferredembodiments, other suitable speed control parameters may be used, asneeded or desired, giving due consideration to the goals of adaptivelycontrolling knee resistances at various speeds, and/or of achieving oneor more of the benefits and advantages as taught or suggested herein.

The foot contact time is preferably measured or computed during aparticular time period. Preferably, the foot contact time is measuredduring one stance phase. Alternatively, the foot contact time may bemeasured or computed over one or more gait cycles. The foot contact timeis preferably computed based on signals from the axial force sensors140. A nonzero axial force measurement indicates that the prostheticfoot is in contact with the ground or other supporting surface.

Referring to FIG. 7, typically, as walking speed increases, foot contacttime decreases. In FIG. 7, foot contact time for one subject is plottedagainst forward walking and running speed, showing decreasing times withincreasing speeds. The x-axis 192 represents the forward speed in cm/secand the y-axis 194 represents the foot contact time during one stancephase in seconds.

In FIG. 7, triangles show contact times for a non-amputee moving atseveral distinct steady state speeds from slow walking at 0.85meters/sec to moderate running at 1.74 meters/sec. As seen in FIG. 7,contact time generally decreases with increasing speed. A least-squaresregression line is fitted to the data with a slope of about −0.32sec²/meter. Similar regressions were observed for both amputees andnon-amputees. Data were collected using a four-camera bilateralkinematic data-acquisition system based on Selspot II cameras fromSelective Electronics Co., Partille, Sweden (Unpublished data fromMassachusetts General Hospital Gait Laboratory, Boston, Mass.).

In accordance with one preferred embodiment, the controller 132 throughan iterative process determines how swing phase knee resistances ordamping are modulated with foot contact time or locomotory speed. Thefull biological range of foot contact time is stored in the memory 150of the knee's processor 132. Typically, a person of short stature has,on average, smaller foot contact times compared with a person of tallstature. The full biological range stored in the memory 150 preferablyincludes both these extremes.

In one preferred embodiment, the memory 150 stores a foot contact timeof zero to about two seconds which is generally more than sufficient tocover the full biological range of foot contact times. In otherpreferred embodiments, the memory may store a smaller or larger range offoot contacts times with efficacy, as required or desired, giving dueconsideration to the goals of covering the full biological range of footcontact times, and/or of achieving one or more of the benefits andadvantages as taught or suggested herein.

Preferably, the foot contact time range is partitioned into time slotsor partitions within the microprocessor memory 150. When an amputeemoves from a slow to a fast walk different time slots or locomotoryvelocity ranges are sampled. Since the entire biological range ispartitioned, each amputee, independent of height, weight or body size,samples multiple time slots when moving from a slow to a fast walk orrun.

In one preferred embodiment, the partition size is about 100milliseconds (msecs), thus giving a total of twenty time slots over atwo-second foot contact time range or interval. Any one amputee wouldtypically sample not all but a fraction of the twenty time slots whenmoving from a slow to a fast locomotory pace. In other preferredembodiments, the partition size can be alternately selected withefficacy, as required or desired, giving due consideration to the goalsof achieving one or more of the benefits and advantages as taught orsuggested herein.

The control scheme of one preferred embodiment preferably modulates kneedamping profiles within each time slot. In State 4, damping values aremodulated within each time slot to control peak flexion angle, and inState 5, the impact force of the extending leg against the artificialknee cap is controlled. Based on sensory data provided to the controller132 (as discussed above), the controller 132 sends appropriate commandsignals or instructions to the knee brake or damper 130.

State 4 (Swing Flexion):

When an amputee first walks or takes a first step utilizing theprosthetic knee system 110 as controlled by the preferred controlschemes of the invention, preferably, the microprocessor or controller132 initially sets or adjusts State 4 damping or resistance to kneerotation to its lowest value within each time slot. Hence, when anamputee takes a first step, State 4 knee damping torque is minimized,and the knee swings freely throughout early swing phase.

Preferably, this minimum value of the knee damping torque is about 0.4N-m and is largely determined by the particular knee brake utilized.Alternatively, other minimum damping torque values and/or variabletorques may be utilized with efficacy, as needed or desired, giving dueconsideration to the goals of achieving one or more of the benefits andadvantages as taught or suggested herein.

For subsequent steps or gait cycles, after the first step, thecontroller 132 preferably increases brake damping by sending appropriatecommand signals or instructions to the knee brake 130 whenever the kneeflexes to an angle greater than a fixed or predetermined target angle.For walking non-amputees, peak flexion angle during early swingtypically does not exceed about 80° (see FIG. 3).

Hence, in accordance with one preferred embodiment, to achieve a gaitcycle that is substantially natural or biological, the target angle isset equal to about 80° to control the State 4 peak flexion angle of theprosthetic knee system 10. In other preferred embodiments, and/or otheractivity levels or external conditions, the State 4 target angle can bealternately selected, as needed or desired, giving due consideration tothe goals of providing a substantially life-like response, and/or ofachieving one or more of the benefits and advantages as taught orsuggested herein.

The microprocessor 132 preferably increases damping by an amount that isproportional to the error or difference between the actual flexionangle, measured by the angle sensor 134, and the target angle. Increaseddamping lowers the peak flexion angle for future gait cycles, butpreferably only in those time slots or locomotory speeds which theamputee has sampled.

In State 4, when the peak flexion angle falls below the target angle themicroprocessor 132 decreases the damping torque by sending appropriatecommand signals or instructions to the knee brake 130. This ensures thatdamping levels are not unnecessarily high.

Preferably, the damping torque is decreased when the peak flexion anglefalls below the target angle for N consecutive locomotory steps, cyclesor strides. One preferred value for N is about twenty locomotory or gaitcycles, though other values may be efficaciously utilized. The brakedamping is preferably decreased by an amount proportional to the erroror difference between the actual flexion angle, measured by the anglesensor 134, and the target angle. Within any particular time slot orbin, decreased damping raises the peak flexion angle for future gaitcycles.

Typically, at faster walking speeds, a greater damping level is requiredto keep the peak flexion angle in State 4 below the target anglethreshold. Hence, to increase State 4 adaptation speed, in one preferredembodiment, the control scheme is designed such that damping levels atfaster walking speeds or time slots are at least as high as dampinglevels at slower speeds or time slots.

Moreover, preferably, the State 4 damping levels applied in each timeslot over one gait or locomotory cycle are constant, though they may bevariable or angle dependent. Additionally, the modulation of State 4damping levels in one or more time slots may involve changing thedamping over a fixed or predetermined knee angle range or changing theangle range over which damping is applied or a combination thereof.

As the amputee continues to use the prosthetic knee system 110 andsamples a diverse range of walking, running or other locomotory speeds,State 4 knee damping gradually converges within each time slot untilpeak knee flexion always falls below, or close to, the target angle forsubstantially all walking, running or other locomotory speeds. Theoptimized damping torque values or profiles for each time slot orlocomotory speed are stored in the microprocessor memory 150. Hence,once the iterative adaptive control scheme has been implemented, theamputee can rapidly accelerate from a slow to a fast walk all the whilesampling different time slots, and therefore, different damping levelswithin State 4.

State 5 (Swing Extension):

A similar scheme or strategy is used to control the force of impact whenthe swinging prosthesis strikes the artificial knee cap. As noted above,this artificial knee cap serves as an extension stop.

When an amputee first walks or takes a first step utilizing theprosthetic knee system 110 as controlled by the preferred controlschemes of the invention, preferably, the microprocessor or controller132 initially sets or adjusts State 5 damping to its lowest value withineach time slot. Hence, when an amputee takes a first step, State 5 kneedamping torque is minimized, and the knee extends from the peak flexionangle in State 4 to the maximum extension angle (about 180°) in State 5.Contact with the artificial knee cap prevents further extension.

Preferably, this minimum value of the knee damping torque is about 0.4N-m and is largely determined by the particular knee brake utilized.Alternatively, other minimum damping torque values and/or variabletorques may be utilized with efficacy, as needed or desired, giving dueconsideration to the goals of achieving one or more of the benefits andadvantages as taught or suggested herein.

For subsequent steps or gait cycles, after the first step, thecontroller 132 computes an average impact force of the swinging legagainst the artificial kneecap, within each bin or time slot, with thedamping minimized. From the smallest of the M time slots or bins to thelargest, if two consecutive bins are not directly adjacent then a linearextrapolation is performed to estimate the average impact forces forintermediate bins. For example, if averages are computed for bins “ten”and “twelve”, but not for bin “eleven”, then a linear extrapolation fromthe impact force corresponding to bin “ten” to the impact forcecorresponding to bin “twelve” is computed. This linear function is thenemployed to estimate an impact force for bin “eleven”. The M bin regionpreferably comprises between about three to five bins or time slots,though fewer or more may be efficaciously used, as needed or desired.

After M average impact forces are computed and linear extrapolations areformulated from the minimum to the maximum bins, knee damping values areselected using a clinically determined relationship relating impactforce to optimal extension damping. Hence, the amputee feels dampingtending to decelerate the extending leg but only for walking speedscorresponding to the M bin region. For bins above the maximum and belowthe minimum, the default minimum damping is used until additional dataare collected and average impact forces are computed. For bins above andbelow the original M bin region, linear extrapolations are preformed toestimate average impact forces for intermediate bins. For example, ifthe maximum of the original M bins is equal to “fourteen”, and anaverage impact force is computed for bin “seventeen”, then impact forcesare estimated for bins “fifteen” and “sixteen” using a linear functionfrom the average impact force corresponding to bin “fourteen” and theaverage force corresponding to bin “seventeen”. Once average impactforces are computed for bins above and below the region of the originalM bins, knee damping values are selected using a clinically determinedrelationship relating impact force to optimal extension damping.

The clinically determined relationship relating impact force to optimalextension damping is preferably derived or determined by a clinicalinvestigation utilizing patients moving at different walking, runningand or other locomotory speeds. Preferably, the clinically determinedrelationship relating impact force to optimal extension damping isderived or determined by a clinical investigation utilizing patientshaving different body sizes (weights). This clinically determinedrelationship is preferably stored in the system memory 150.

For each time slot or bin, once an optimal extension damping value hasbeen selected, the microprocessor 132 once again computes an averageimpact force, and this new average force is then used as a target. If asystem disturbance occurs that significantly alters the magnitude ofimpact force within a particular bin, then extension damping ismodulated until the impact force is once again equal to, or in theproximity of, the target impact force. For example, within a particularbin, if the average impact force after the damping is turned on is 100Newtons, and a disturbance causes the swinging leg to impact theartificial kneecap with a force of 150 Newtons, then extension dampingis increased for that bin until the impact force is once again equal to,or approximately close to, the original 100 Newtons. With this adaptiveroutine, the amputee can change from a lightweight shoe to a heavy shoeand still walk comfortably without having to return to their prosthetistfor re-programming.

The average impact force of the swinging leg against the artificialkneecap is preferably computed by the controller 132 using signals ordata provided by sensors local to the prosthesis. The impact forcesensors preferably comprise the sensors 140 and include one or morestrain gauges mounted on or mechanically connected to the frame 141, asdiscussed above. Based on the computed or determined impact force, thecontroller 132 provides appropriate command signals or instructions tothe knee brake 130 to control the knee damping.

State 5 damping, in each time slot or locomotory speed, can be modulatedby several methods in the preferred embodiments of the control scheme ofthe invention. For example, the modulation of State 5 damping levels inone or more time slots may involve changing the damping over a fixed orpredetermined knee angle range or changing the angle range over whichdamping is applied or a combination thereof. Additionally, State 5damping levels applied in one or more time slots over one gait orlocomotory cycle may be constant, variable and/or angle dependent.

In accordance with one preferred embodiment, the control schememodulates the knee damping in State 5 over or within a fixed orpredetermined angle range. For example, knee damping torque is increasedor decreased within a particular extension angle range such as in therange from about 130° to about 180° to increase or decrease the dampingwithin that particular time slot.

In accordance with another preferred embodiment, the control schemekeeps the State 5 knee damping levels substantially constant and insteadmodulates the angle range over which knee damping is applied. Forexample, the knee damping is constant and maximized, and this damping isapplied over an extension angle range of about 170° to about 180°. Toincrease State 5 damping, the starting extension angle for theinitiation of knee damping could be changed from about 170° to about160° to increase the State 5 damping for that particular time slot orlocomotory speed.

Typically, at faster walking speeds, a greater damping level is requiredto keep the impact force against the artificial kneecap at an acceptablerange. Hence, to increase State 5 adaptation speed, in one preferredembodiment, the control scheme is designed such that damping levels atfaster walking speeds or time slots are at least as high as dampinglevels at slower speeds or time slots.

As the amputee continues to use the prosthetic knee system 110 andsamples a diverse range of walking and running speeds, State 5 kneedamping gradually converges within each time slot until the impactforces of the swinging leg against the artificial kneecap are held at anacceptable level for substantially all walking, running or otherlocomotory speeds. The optimized damping torque values or profiles foreach time slot or locomotory speed are stored in the microprocessormemory 150. Hence, once the iterative adaptive control scheme has beenimplemented, the amputee can rapidly accelerate from a slow to a fastwalk all the while sampling different time slots, and therefore,different damping levels within State 5.

As the patient further continues to use the prosthetic knee system 110,further automated refinements and fine-tuning can be made by the system110, as necessary. The prosthesis of the preferred embodiments is aself-teaching and/or self-learning system that is guided by clinical(prosthetic) and biomechanical knowledge. For example, biomechanicalknowledge (stored in the system memory) includes information related tothe mechanics of typical human walking/running, as discussed above inreference to FIG. 1.

Advantageously, no patient-specific is needed by the control scheme andprosthetic knee system of the preferred embodiments, and hence nopre-programming by a prosthetist or amputee is needed to accommodatedifferent locomotory speeds and different patients. The system is ableto adapt to various types of disturbances once the patient leaves theprosthetist's facility because it is patient-adaptive andspeed-adaptive. Desirably, this also saves on time and cost, andsubstantially eliminates or mitigates inconvenience, discomfort andfatigue for the patient during an otherwise lengthy adjustment or trialperiod.

The control scheme and prosthesis of the preferred embodiments allow thepatient to perform a wide variety of activities. These include normalwalking or running on a level or inclined surface, sitting down,ascending or descending steps or other situations, for example, when auser lifts a suitcase.

Magnetorheological Knee Brake

Preferred embodiments of a magnetorheological knee brake or actuator inaccordance with the present invention are described in copending U.S.application Ser. No. 09/767,367, filed Jan. 22, 2001, entitled“ELECTRONICALLY CONTROLLED PROSTHETIC KNEE,” the entire disclosure ofwhich is hereby incorporated by reference herein. For purposes ofclarity and brevity of disclosure, only a brief description of thismagnetorheological knee brake or actuator is set forth below.

FIG. 8 is a simplified schematic of a rotary prosthetic knee brake ormagnetorheological (MR) braking system 210 in accordance with onepreferred embodiment of the present invention. The knee actuator 210includes a substantially central core 212 substantially circumscribed orenveloped by an electromagnet or magnetic coil 214 and in mechanicalcommunication with a pair of side plates or disks 216, 218. By passing avariable, controlled current through the electromagnet 214, a variablemagnetic field is created. Preferably, the core 212 and side plates 216,218 are fabricated from a ferrous, magnetizable or magnetic material andthe like. More preferably, the core 212 and side plates 216, 218 arefabricated from a magnetically soft material of high flux saturationdensity and high magnetic permeability.

The prosthetic knee brake or actuator 210 further includes a pluralityof inner blades or plates 220 in mechanical communication with an innerspline 222. The inner spline 222 generally circumscribes or envelops theelectromagnet 214 and is coupled or mechanically connected to the sideplates 216, 218. The blades 220 are preferably concentrically arrangedabout the brake axis of rotation 224. The inner spline 222 is preferablyrotatable about the knee joint axis of rotation 224, and hence so arethe blades or rotors 220 and the core side plates 216, 218. Rotation ofthe inner spline 222 corresponds to rotation or movement of the lower(below the knee) part of the leg.

The prosthetic knee brake or actuator 210 also comprises a plurality ofouter blades or plates 230 in mechanical communication with an outerspline 232. The outer spline 232 generally circumscribes or envelops theinner spline 222. The blades 230 are preferably concentrically arrangedabout the brake axis of rotation 224. The outer spline 232 is preferablyrotatable about the knee joint axis of rotation 224, and hence so arethe blades or stators 230. Rotation of the outer spline 232 correspondsto rotation or movement of the upper (above the knee) part of the leg.Preferably, the outer spline or housing 232 comprises means tofacilitate connection of the prosthetic knee joint 210 to a suitablestump socket or the like. The outer spline 232, and hence the stators230, are preferably substantially irrotationally coupled to ornonrotatable with respect to the stump socket or residual limb.

The plurality of rotors 220 and stators 230 are interspersed in analternating fashion and the gaps between adjacent blades 220 and 230comprise a magnetorheological (MR) fluid 234, which thereby resides inthe cavity or passage formed between the inner spline 222 and the outerspline 232. In one preferred embodiment, the MR fluid 234 in the gaps ormicrogaps between adjacent rotors 220 and stators 230 is in the form ofthin lubricating films between adjacent rotors 220 and stators 230.Shearing of MR fluid present between the side plates 216, 218 andadjacent stators 230 can also contribute to the knee damping.

During knee joint rotation, the MR fluid in the plurality of gapsbetween the rotors 220 and stators 230 is sheared to generate a dampingtorque to control the limb rotation. The blades or disks 220 and 230 arepreferably formed of a ferrous, magnetizable or magnetic material andthe like. More preferably, the blades or disks 220 and 230 are formed ofa material of as high magnetic permeability and magnetic softness as ismechanically practical.

The knee joint actuator 210 further includes a pair of ball bearings226, 228 coupled or connected to the respective side plates 216, 218.The ball bearings 226, 228 are further coupled or connected torespective side walls or mounting forks 236, 238. Thus, a rotarycoupling is created between the inner spline 222 and the mounting forks236, 238. The mounting forks 236, 238 in combination with the outerspline 232 form one main outer shell of the knee actuator 210.Preferably, the side walls or mounting forks 236, 238 comprise means tofacilitate connection of the prosthetic knee actuator 210 to a suitablepylon, shank portion or the like.

Preferably, the central core 212 and the electromagnet 214 also rotatealong with the rotation of the inner spline 222, the rotors 220, thecore side plates 216, 218 and the mounting forks 236, 238. The stators230 rotate together with the rotation of the outer spline 232.

The rotors 220 are rotationally fixed relative to the inner spline 222and the stators 230 are rotationally fixed relative to the outer spline232. During various stages of locomotion or knee rotation, and about theknee axis of rotation 224, the rotors 220 may rotate while the stators230 are rotationally substantially stationary, or the stators 230 mayrotate while the rotors 220 are rotationally substantially stationary,or both the rotors 220 and the stators 230 may rotate or besubstantially rotationally stationary. The terms “rotor” and “stator”are used to distinguish the inner blades 220 and the outer blades 230,though both rotors 220 and stators 230 can rotate, and teach thatrelative rotational motion is created between the rotors 220 and thestators 230 (with MR fluid being sheared in the gaps between adjacentrotors 220 and stators 230). If desired, the blades 220 can be referredto as the “inner rotors” and the blades 230 as the “outer rotors.”

Actuation of the magnet 214 causes a magnetic field, circuit or path 240to be generated or created within the knee actuator 210. In onepreferred embodiment, the magnetic field 240 passes through the centralcore 212, radially outwards through the side plate 218, laterallythrough the interspersed set of rotors 220 and stators 230 and themagnetorheological fluid 234, and radially inwards through the sideplate 216. The portion of the magnetic field 240 passing through thecore 212 and side plates 216, 218 generally defines the magnetic returnpath while the active or functional magnetic field is generally definedby the magnetic path through the rotors 220, stators 230 and MR fluid234.

The magnetorheological (MR) fluid 234 undergoes a rheology or viscositychange which is dependent on the magnitude of the applied magneticfield. In turn, this variation in fluid viscosity determines themagnitude of the shearing force/stress, torque or torsional resistancegenerated, and hence the level of damping provided by the prostheticknee brake 210. Thus, by controlling the magnitude of this magneticfield, the rotary motion of the artificial limb is controlled, forexample, to control the flexion and extension during swing and stancephases to provide a more natural and safe ambulation for the amputee.

In one preferred embodiment, the rotors 220 and/or stators 230 aredisplaceable in the lateral direction 242, and hence under the influenceof a magnetic field can rub against adjacent rotors 220 and/or stators230 with a variable force determined by the strength of the magneticfield to create a “hybrid” magnetorheological and frictional dampingbrake. In another preferred embodiment, the rotors 220 and stators 230are laterally fixed in position relative to the splines 222 and 232, andhence the braking effect is substantially purely magnetorheological orviscous. Alternatively, some of the rotors 220 and/or stators 230 may belaterally fixed while others may be laterally displaceable, as requiredor desired, giving due consideration to the goals of providing asubstantially natural feeling and/or safe prosthetic device, and/or ofachieving one or more of the benefits and advantages as taught orsuggested herein. In one embodiment, the side plates 216, 218 arelaterally displaceable and contribute to the frictional damping due tofrictional contact with adjacent stators 230.

Advantageously, by operating in the shear mode, there is no ornegligible pressure build-up within the MR actuated prosthetic knee ofthe present invention. This substantially eliminates or reduces thechances of fluid leakage and failure of the knee, and hence desirablyadds to the safety of the device.

Also advantageously, the multiple shearing surfaces or flux interfaces,provided by the preferred embodiments of the present invention, behavelike a torque multiplier and allow the viscous torque level to bestepped up to a desired maximum value without the use of an additionaltransmission or other auxiliary component. For example, if two fluxinterfaces can provide a maximum viscous torque of about 1 N/m, thenforty flux interfaces will be able to provide a viscous damping torqueof about 40 N/m. In contrast, if a 40:1 step-up transmission is used toincrease the viscous torque, disadvantageously, not only is the systemreflected inertia magnified by a factor of about 1600, but the systemweight, size and complexity are undesirably increased.

The multiple shearing surfaces or interfaces of the prosthetic kneeactuator of the preferred embodiments also advantageously allow for awide dynamic torque range to be achieved which permits safe and/or morenatural ambulation for the patient. Desirably, the MR actuatedprosthetic knee of the preferred embodiments provides a rapid andprecise response. Again, this permits the patient to move in a safeand/or more natural manner.

FIGS. 9 and 10 show a magnetorheological rotary prosthetic kneeactuator, brake or damper 210 having features and advantages inaccordance with one preferred embodiment of the present invention. Theprosthetic knee actuator 210 generates controllable dissipative forcespreferably substantially along or about the knee axis of rotation 224.The knee actuator embodiment of FIGS. 9 and 10 is generally similar inoperation and structure to the knee actuator embodiment of FIG. 8, andhence for purposes of clarity and brevity of disclosure only a briefdescription of the embodiment of FIGS. 9 and 10 is set forth below.

The electronically controlled knee actuator 210 generally comprises agenerally central core 212 in mechanical communication with a pair ofrotatable side plates 216, 218, an electromagnet 214, a plurality ofblades or rotors 220 in mechanical communication with a rotatable innerspline 222, a plurality of blades or stators 230 in mechanicalcommunication with a rotatable outer spline 232, a pair of ball bearings226, 228 for transferring rotary motion to a pair of outer side walls orforks 236, 238. The rotation is substantially about the knee axis ofrotation 224.

The plurality of rotors 220 and stators 230 are preferably interspersedin an alternating fashion and the gaps or microgaps between adjacentblades 220 and 230 comprise thin lubricating films of amagnetorheological (MR) fluid, which thereby resides in the cavity orpassage formed between the inner spline 222 and the outer spline 232.This preferred embodiment provides a controllable and reliableartificial knee joint, which advantageously has a wide dynamic torquerange, by shearing the MR fluid in the multiple gaps or flux interfacesbetween adjacent rotors 220 and stators 230.

Preferably, end-threaded rods 248 and nuts 250 are used to secureselected components of the prosthetic knee 210, thereby allowing astraightforward assembly and disassembly procedure with a minimum offasteners. Alternatively, or in addition, various other types offasteners, for example, screws, pins, locks, clamps and the like, may beefficaciously utilized, as required or desired, giving due considerationto the goals of providing secure attachment, and/or of achieving one ormore of the benefits and advantages as taught or suggested herein.

In one preferred embodiment, the prosthetic knee brake 210 furthercomprises a flexion stop system or assembly. The flexion stop systemcontrols the maximum allowable flexion angle by physically limiting therotation between the outer side forks 236, 238 and the outer spline 232,and hence the rotation of the knee joint.

In one preferred embodiment, the prosthetic knee brake 210 furthercomprises an extension stop system or assembly. The extension stopsystem controls the maximum allowable extension angle by physicallylimiting the rotation between the outer side forks 236, 238 and theouter spline 232, and hence the rotation of the knee joint.

In one preferred embodiment, the prosthetic knee brake 210 furthercomprises an extension assist to help straighten the leg by urging orbiasing the leg to extension by applying a controlled torque or force.Any one of a number of devices, such as a spring-loaded extensionassist, as known in the art may be used in conjunction with the presentinvention.

In one preferred embodiment, the prosthetic knee brake 210 comprisesforty rotors 220 and forty one stators 230 interspersed in analternating fashion. This results in forty flux interfaces or fluid gapsin which the magnetorheological (MR) fluid resides. In another preferredembodiment, the number of rotors 220 is about ten to one hundred, thenumber of stators 230 is about eleven to one hundred one so that thenumber of MR fluid to rotor interfaces which produce braking in thepresence of a magnetic field is twice the number of rotors. In yetanother preferred embodiment, the number of rotors 220 is in the rangeof one to one hundred. In a further preferred embodiment, the number ofstators 230 is in the range of one to one hundred. In other preferredembodiments, the number of rotors 220, stators 230 and/or fluxinterfaces may be alternately selected with efficacy, as needed ordesired, giving due consideration to the goals of providing a widedynamic torque range, and/or of achieving one or more of the benefitsand advantages as taught or suggested herein.

Advantageously, the induced yield stress or viscous torque isproportional to the overlap area between a rotor-stator pair multipliedby twice the number of rotors (the number of MR fluid to rotorinterfaces which produce braking torque in the presence of a magneticfield). This desirably allows the viscous torque or yield stress to beincreased or decreased by selecting or predetermining the number ofrotors 220 and/or stators 230 and/or the overlap or mating surface areabetween adjacent rotors 220 and/or stators 230. Another advantage isthat this permits control over the overall size, that is radial size andlateral size, of the MR actuated prosthetic brake 210. For example, theoverall knee configuration may be made radially larger and laterallyslimmer while providing the same viscous torque range by appropriateselection of the number of flux interfaces and the overlap area of theshearing surfaces.

It is desirable to minimize the MR fluid gap between adjacent rotors 220and stators 230 since the power needed to saturate the total MR fluidgap is a strong function of the gap size. Thus, advantageously, asmaller gap size renders the MR actuated brake 210 more efficient andreduces power consumption.

Preferably, the MR fluid gap size is also selected so that in theabsence of an applied magnetic field only a viscous damping force ortorque component is present from the shearing of MR fluid betweenadjacent rotor and stator surfaces. That is, there is no frictionaltorque component between the rotors 220 and stators 230 under zero-fieldconditions.

Accordingly, in one preferred embodiment, the power required to saturatethe MR fluid is lowered and the dynamic range of the knee is enhanced byminimizing the MR fluid gap size. In this embodiment, the gap is notreduced so much that, under zero-field conditions, a normal force actsbetween adjacent rotor and stator surfaces, causing frictional rubbing.The absence of friction between rotors and stators enables the kneejoint to swing freely, thereby providing a wider dynamic range. As anote, the viscous damping at zero-field does not increase dramaticallywith decreasing fluid gap because the MR fluid exhibits a property knownas shear rate thinning in which fluid viscosity decreases withincreasing shear rate.

In one preferred embodiment, the MR fluid gap size or width betweenadjacent rotors 220 and stators 230 is about 40 microns (μm) or less. Inanother preferred embodiment, the MR fluid gap size or width betweenadjacent rotors 220 and stators 230 is in the range from about 10 μm toabout 100 μm. In other preferred embodiments, the MR fluid gap size canbe alternately dimensioned and/or configured with efficacy, as requiredor desired, giving due consideration to the goals of providing an energyefficient prosthetic knee actuator 210 having a wide dynamic torquerange, and/or of achieving one or more of the benefits and advantages astaught or suggested herein.

The electronically controlled magnetorheologically actuated prostheticknee brake of the preferred embodiments provides high-speed instantlyresponsive control of knee movement, yet is robust and affordable forthe amputee. The preferred embodiments advantageously provide improvedstability, gait balance and energy efficiency for amputees and simulateand/or closely recreate the dynamics of a natural knee joint.

During operation, the electromagnet or magnetic coil 214 is actuated, asneeded, by a selected or predetermined electrical signal, voltage orcurrent to generate an active variable magnetic field passingsubstantially perpendicularly to the plurality of rotor and statorsurfaces and through the MR fluid or film between adjacent rotors 220and stators 230 to generate a variable damping torque (or rotaryresistive force) which precisely and accurately controls the rotarymotion of the prosthetic knee 210. As discussed above, in accordancewith one preferred embodiment, the torque comprises a frictional dampingcomponent.

Desirably, the MR actuated prosthetic knee 210 of the preferredembodiments provides a rapid and precise response. The materials in MRparticles respond to the applied magnetic field within milliseconds,thereby allowing for real-time control of the fluid rheology and theknee motion. This facilitates in permitting the patient to move in asafe and/or more natural manner.

Advantageously, the viscous damping torque is generated by shearing ofthe MR fluid. Hence, there is no or negligible pressure build-up orchange within the MR actuated prosthetic knee 210 of the presentinvention. This substantially eliminates or reduces the chances of fluidleakage and failure of the knee, and hence desirably adds to the safety.Moreover, costly and/or relatively complex components such as pressurebearings and the like need not be utilized to provide a reliable seal.

Another advantage is that the plurality of shearing surfaces or fluxinterfaces between adjacent rotors 220 and stators 230 behave like atorque multiplier and allow the viscous torque level (and/or frictionaltorque) to be stepped up to a desired maximum value without the use ofan additional transmission or other auxiliary component. Moreover, theflexibility in selecting the overlap surface area between adjacentrotors 220 and stators 230 can also increase or decrease the maximumattainable viscous torque (and/or frictional torque). Thus, desirably awide dynamic torque or torsional resistance range can be provided, asneeded or desired, which adds to the versatility of the inventionwithout adding substantially to system size, weight and complexity.

In one preferred embodiment, the prosthetic knee actuator of thepreferred embodiments provides a maximum dynamic torque of about 40Newton-meters (N-m). In another preferred embodiment, the prostheticknee actuator of the preferred embodiments provides a dynamic torque inthe range from about 0.5 N-m to about 40 N-m. In yet another preferredembodiment, the prosthetic knee actuator of the preferred embodimentsprovides a dynamic torque in the range from about 1 N-m to about 50 N-m.In other preferred embodiments, the prosthetic knee actuator can provideother dynamic torque ranges with efficacy, as needed or desired, givingdue consideration to the goals of achieving one or more of the benefitsand advantages as taught or suggested herein.

Also advantageously, the optimized thinness of the MR fluid gap betweenadjacent rotors 220 and stators 230 provides a higher maximum torque, awider dynamic torque range and requires less energy consumption,preferably about 10 Watts or less. This adds to the efficiency andpracticality of the MR actuated prosthetic brake 210 of the preferredembodiments and also saves on cost since a lower wattage and/or lesscomplex power source can be used.

While the components and techniques of the present invention have beendescribed with a certain degree of particularity, it is manifest thatmany changes may be made in the specific designs, constructions andmethodology hereinabove described without departing from the spirit andscope of this disclosure. It should be understood that the invention isnot limited to the embodiments set forth herein for purposes ofexemplification, but is to be defined only by a fair reading of theappended claims, including the full range of equivalency to which eachelement thereof is entitled.

1. A method of adaptively controlling a prosthetic joint worn by anamputee, comprising: measuring speed indicative data with sensors localto said prosthetic joint as said amputee moves at various speeds;storing said data in a memory of said prosthetic joint in binscorresponding to the speed of said amputee; iteratively modulating thedamping to achieve a predetermined and/or computed target until thedamping converges within each bin; and controlling said prosthetic jointby utilizing the converged damping values to control the damping of saidprosthetic joint by varying the viscosity of a magnetorheological fluidcontained in said prosthetic joint that provides variable resistance toflexion and/or extension, wherein said damping is created primarily byshear forces, wherein said prosthetic joint comprises a prosthetic knee,and wherein said method further comprises controlling said prostheticknee worn by said amputee using a controller which transitions between aplurality of states of biological gait including a first state generallycorresponding to stance flexion, a second state generally correspondingto stance extension, a third state generally corresponding to kneebreak, a fourth state generally corresponding to swing flexion, and afifth state generally corresponding to swing extension, said methodfurther comprising: measuring sensory information and providing saidsensory information to said controller for computation of axial force,extension moment, knee angle and velocity; transitioning from said firststate to said second state under condition C12 and said condition C12being satisfied when said prosthetic knee achieves a predeterminedextension velocity; transitioning from said second state to said thirdstate under condition C23 and said condition C23 being satisfied whensaid extension moment is below a first threshold; transitioning fromsaid third state to said fourth state under condition C34 and saidcondition C34 being satisfied when said axial force falls below a secondthreshold; transitioning from said fourth state to said fifth stateunder condition C45 and said condition C45 being satisfied when saidprosthetic knee begins to extend; transitioning from said fifth state tosaid first state under condition C51 and said condition C51 beingsatisfied when said axial force climbs above a third threshold; andcontrolling operation of said prosthetic knee in said states ofbiological gait by processing said sensory information to provide acontrolled and variable resistance to flexion and/or extension.
 2. Themethod of claim 1, wherein measuring speed indicative data comprisesmeasuring ground contact time.
 3. The method of claim 2, wherein saidground contact time changes substantially monotonically with the speedof said amputee.
 4. The method of claim 2, wherein said ground contacttime is stored in about twenty said bins.
 5. The method of claim 4,wherein each of said bins has a size that represents about 100milliseconds.
 6. The method of claim 1, wherein iteratively modulatingthe damping comprises iteratively modulating the damping to achieve atarget swing flexion angle over a range of speeds of the amputee.
 7. Themethod of claim 6, wherein said target swing flexion angle is about 80°.8. The method of claim 1, wherein controlling said prosthetic jointcomprises controlling flexion and/or extension of said prosthetic jointover a range of speeds of the amputee.
 9. The method of claim 1, whereinsaid method further comprises applying a magnetic field to vary theviscosity of said magnetorheological fluid.
 10. The method of claim 9,wherein said method further comprises shearing said magnetorheologicalfluid.
 11. The method of claim 10, wherein shearing saidmagnetorheological fluid comprises shearing said magnetorheologicalfluid in a plurality of gaps.
 12. The method of claim 10, whereinshearing said magnetorheological fluid comprises shearing saidmagnetorheological fluid in a plurality of gaps formed between aplurality of blades.
 13. The method of claim 12, wherein said blades arerotatable.
 14. The method of claim 13, wherein said blades are inmechanical communication with a lower leg portion.
 15. The method ofclaim 13, wherein said blades are in mechanical communication with anupper leg portion.
 16. The method of claim 1, wherein said condition C23is further satisfied when said prosthetic knee is at or close to fullextension.
 17. The method of claim 16, wherein said condition C23 isfurther satisfied when said prosthetic knee has been substantially stillfor a predetermined time.
 18. The method of claim 1, wherein said methodfurther comprises transitioning from said first state to said thirdstate under condition C13 and said condition C13 is satisfied when saidextension moment is below a fourth threshold.
 19. The method of claim18, wherein said condition C13 is further satisfied when said prostheticknee is at or close to full extension.
 20. The method of claim 19,wherein said condition C13 is further satisfied when said prostheticknee has been substantially still for a predetermined time.
 21. Themethod of claim 1, wherein said method further comprises transitioningfrom said first state to said fourth state under condition C14 and saidcondition C14 is satisfied when said axial force falls below a fourththreshold.
 22. The method of claim 1, wherein said method furthercomprises transitioning from said second state to said first state undercondition C21 and said condition C21 is satisfied when said prostheticknee achieves a predetermined flexion velocity.
 23. The method of claim1, wherein said method further comprises transitioning from said secondstate to said fourth state under condition C24 and said condition C24 issatisfied when said axial force falls below a fourth threshold.
 24. Themethod of claim 1, wherein said method further comprises transitioningfrom said third state to said first state under condition C31 and saidcondition C31 is satisfied when said prosthetic knee has been in saidthird state for a predetermined time.
 25. The method of claim 1, whereinsaid method further comprises transitioning from said third state tosaid first state under condition C31 and said condition C31 is satisfiedwhen extension moment is above a fourth threshold.
 26. The method ofclaim 25, wherein said condition C31 is further satisfied when saidprosthetic knee is at or close to full extension.
 27. The method ofclaim 1, wherein said method further comprises transitioning from saidfourth state to said first state under condition C41 and said conditionC41 is satisfied when said axial force climbs above a fourth threshold.28. The method of claim 1, wherein said prosthetic knee comprises amagnetorheological damper.